The following relates to the imaging, diagnostic, and related arts. It finds particular application in simultaneous magnetic resonance (MR) imaging and positron emission tomography (PET) imaging, and is described with particular reference thereto. However, the following finds more general application in acquisition of PET and MR data from a common subject performed simultaneously, sequentially, in a time-interleaved fashion, or by some combination thereof, and to diagnostic processes using same such as imaging, magnetic resonance spectroscopy, and so forth.
MR and PET imaging are imaging modalities that can sometimes provide more information operating in concert than is provided by either modality operating alone. To maximize the synergy of combined MR and PET imaging, it would be useful to perform simultaneous MR and PET imaging, or at least to perform MR and PET imaging together over a relatively short time interval, for example while the subject remains stationary on a common patient bed with respect to the short time interval. Such integrated data acquisition would simplify spatial and optional temporal registration of images acquired by MR and PET, and would reduce the likelihood of occurrence of an operatively significant change in the patient or other subject between the MR and PET imaging data acquisitions. Other advantages of combined PET/MR imaging include the ability to use MR to construct an attenuation map for use in PET imaging, and the use of MR and PET together for motion compensation.
However, construction of a combined PET/MR scanner (sometimes called a hybrid PET/MR scanner) has heretofore been hindered by detrimental effect of the magnetic field of the MR acquisition sub-system on the photomultiplier tube (PMT) detectors of the PET acquisition sub-system. PMT detectors operate by avalanche multiplication of electrons. In a typical PMT arrangement, a photocathode is biased negatively. A photon striking the photocathode generates an initial burst of one or more electrons that travel through vacuum to a first dynode, where they induce generation of a larger number of electrons that travel through vacuum to a second dynode, and so forth, until the avalanche-multiplied electron burst reaches the anode. For PET applications, the PMT detector is typically arranged to view a scintillator that generates a burst of light (i.e., ultraviolet, visible, and/or infrared photons) responsive to interaction with a 511 keV gamma photon generated by a positron-electron annihilation event.
As the PMT operation is based on travel of electrons (which are charged particles) through a vacuum, a force proportional to the cross-product of the electron charge times the electron velocity and the magnetic field is exerted. This force can be represented, for example, in a form such as F=qv×B where q is the electron charge, v is the electron velocity vector, B is the magnetic field vector, and F is the force exerted on the electron traveling at velocity v by the magnetic field B. The effect of magnetic field on the PMT operation is suitably explained as follows. The process of electron multiplication is done via an electrostatic field E in the PMT. If there is also a static magnetic field B present, the force on the electron is given by F=q(E+v×B) and so the electron no longer accelerates along the direction of the accelerating electric field E for |v|>0. In consequence the acceleration of the electron and thus the multiplication is disturbed by the presence of the magnetic field B. For example, if the velocity of the electron is calculated as a function of time assuming a magnetic field orthogonal to the accelerating electric field, the velocity of the electron is seen to be zero at specific points in time, which corresponds to a restart of the acceleration. The extreme sensitivity of typical PMT detectors is such that even the earth's magnetic field (typically about 3×10−5 T to 6×10−5 T) is sufficient to degrade PMT operation. The small magnetic field of the earth can be compensated by suitable calibration of the PMT detector. In contrast, a typical MR scanner generates a static (B0) magnetic field of 0.2 T to 7 T, depending upon the strength of the main magnet, with higher-field MR scanners in development. The effect of this much larger B0 magnetic field on the PMT operation cannot be adequately compensated by calibration.
One approach for overcoming this problem is to use a solid state radiation detector that is less sensitive to the magnetic field compared with a PMT-based detector. For example, International Publication WO 2006/111869 discloses PET/MR scanners that employ solid state silicon photomultipliers (SiPM) or avalanche photodiodes (APD) as radiation detectors. This solution involves use of SiPM, APD, or other non-conventional detectors.
Another approach for overcoming the PMT/MR magnetic field interaction is to move the PMT detectors out of the MR acquisition system, and hence away from the strong magnetic field. The scintillators remain in or near the MR acquisition system, and are coupled with the remote PMT detectors using fiber optical connections. This approach retains the conventional PMT-based detectors, but has the disadvantages of increased system complexity. An additional disadvantage for time-of-flight (TOF) PET imaging is that the fiber optical connections add significant light loss and transit time spread that adversely affect TOF calculations. Wavelength-dependent dispersion further complicates these TOF calculations and limits the temporal resolution achievable using fiber optical connections.
The following provides new and improved apparatuses and methods which overcome the above-referenced problems and others.